Methods and systems for beam intensity-modulation to facilitate rapid radiation therapies

ABSTRACT

Methods and system for facilitating rapid radiation treatments are provided herein and relate in particular to radiation generation and delivery, electron source design, beam control and shaping/intensity-modulation. The methods and systems described herein are particularly advantageous when used with a compact high-gradient, very high energy electron (VHEE) accelerator and delivery system (and related processes) capable of treating patients from multiple beam directions with great speed, using all-electromagnetic or radiofrequency deflection steering is provided; or when used with a high-current electron accelerator system of energy range more conventionally used in photon radiation therapy to produce much faster delivery of intensity-modulated photon radiation therapy, that can in both cases deliver an entire dose or fraction of high-dose radiation therapy sufficiently fast to freeze physiologic motion, yet with an equal or better degree of dose conformity or sculpting compared to conventional photon therapy.

CROSS-REFERENCES TO RELATED APPLICATIONS

The present application is a Divisional of U.S. Ser. No. 15/068,387filed Mar. 11, 2016 (Allowed); which is a continuation ofPCT/US2014/055252 filed Sep. 11, 2014; which claims priority to U.S.Provisional Application No. 61/876,679 filed Sep. 11, 2013; the entirecontents of which are incorporated herein by reference in their entiretyfor all purposes.

This application is generally related to U.S. application Ser. No.13/765,017 entitled “Pluridirectional Very High Electron EnergyRadiation Therapy Systems and Processes,” filed Feb. 12, 2013 (now U.S.Pat. No. 8,618,521); PCT Application No. PCT/US2014/055260 filed Sep.11, 2014; and PCT Application No. PCT/US2014/055270 filed Sep. 11, 2014;the full disclosures of which are incorporated herein by reference intheir entirety for all purposes.

FIELD OF THE INVENTION

The invention generally relates to radiation therapies and moreparticularly to systems and methods for very rapid radiation therapies.

BACKGROUND OF THE INVENTION

Major technical advances in radiation therapy in the past two decadeshave provided effective sculpting of 3-D dose distributions andspatially accurate dose delivery by imaging verification. Thesetechnologies, including intensity modulated radiation therapy (IMIRT),hadron therapy, and image guided radiation therapy (IGRT) havetranslated clinically to decreased normal tissue toxicity for the sametumor control, and more recently, focused dose intensification toachieve high local control without increased toxicity, as instereotactic ablative radiotherapy (SABR) and stereotactic bodyradiotherapy (SBRT).

One key remaining barrier to precise, accurate, highly conformalradiation therapy is patient, target and organ motion from many sourcesincluding musculoskeletal, breathing, cardiac, organ filling,peristalsis, etc. that occurs during treatment delivery, currently 15-90minutes per fraction for state-of-the-art high-dose radiotherapy. Assuch, significant effort has been devoted to developing “motionmanagement” strategies, e.g., complex immobilization, markerimplantation, respiratory gating, and dynamic tumor tracking.

BRIEF SUMMARY OF THE INVENTION

The present invention relates to methods and systems for facilitatingradiation therapies, particularly extremely rapid radiation therapiesthat rapidly deliver a radiation treatment sufficiently fast enough tofreeze physiologic motion.

In one aspect, the invention relates to a method for treating a patient,that includes: generating one or more patterned particle beams, eachpatterned particle, wherein each of the one or more patterned particlebeams covers an area of the targeted tissue with spatially varying beamintensity according to a treatment pattern of desired radiation dosedistribution; accelerating the one or more patterned particle beams withone or more accelerators; and transporting and/or magnifying thepatterned beam to a desired location, direction, and size suitable forcoverage of the targeted tissue, wherein a shape, resolution andcontrast of the pattern is suitably maintained during transport and/ormagnification so as to deliver the desired radiation dose distributionto the targeted tissue according to the treatment pattern. Magnifyingthe patterned beam may include magnifying the beam through one or morefocusing elements disposed within the beamline of the one or more beams,such as by 100 to 200 times the original size of the pattern. Thefocusing elements may include electro-magnetic lenses, one or morepermanent magnets, electromagnets or a combination thereof. Such methodsmay further include steering the one or more patterned particle beams tothe targeted tissue with one or more beam steering devices from one ormore directions. In some embodiments, steering is concurrent withmagnifying of the one or more patterned particle beams. Methods mayfurther including forming the two-dimensional intensity-modulatedelectron pattern on a photocathode by projecting or scanning a lightsource onto the photocathode.

In certain aspects, the one or more patterned particle beams comprise anarray of smaller patterned beams produced by raster scanning individualsmaller patterned beams from each beam direction of the one or moredirections.

In another aspect, the invention relates to systems for treating apatient, that include: one or more beam generation devices configured togenerate one or more patterned particle beams, each of the one or moreparticle beams covering an area of the targeted tissue with spatiallyvarying beam intensity according to a treatment pattern of desiredradiation dose distribution; one or more accelerators configured foraccelerating the one or more patterned particle beams; and one or moremagnification lenses along a beam line of the one or more particle beamsbetween the accelerators and targeted tissue for magnification of thepatterned particle beam to a desired size suitable for coverage of thetargeted tissue according to the treatment pattern. The system mayinclude one or more beam steering devices configured for steering theone or more patterned particle beams to the targeted tissue from one ormore directions, which may be before or concurrent with magnifying thebeam with one or more lenses, such as with a plurality of small aperturemagnetic lenses.

In some embodiments, the system includes a beam deflector disposed alongbeamlines of the one or more particle beams between the one or moreaccelerators and the targeted tissue such that the one or more beamlines can be directed to the targeted tissue from multiple differingangles using a single common accelerator. The system may be configuredto generate the one or more patterned particle beams from an array ofsmaller patterned beams by raster scanning. The system may be configuredsuch that raster scanning occurs prior to the accelerator while the beamhas a low energy. Raster scanning may be performed by deflection of afixed position electron source or by rastering a laser spot on aphotocathode.

In certain embodiments, the system includes an RF powered or DC particlegun and a photo-cathode configured to produce the two-dimensionalintensity-modulated electron pattern. The system may include aprogrammable controller configured to control deliver the one or moreparticle beams to the targeted tissue from the one or more directionsthereby irradiating the targeted tissue to deliver an entire treatmentdose in less than 10 seconds, preferably about one second or less. Thecontroller is configured to rapidly switch a modulation pattern sent tothe photocathode within a rate of one pattern every 2 seconds or higherso as to provide delivery of differing treatment patterns to thetargeted tissue from multiple directions within less than 10 seconds orless. In one aspect, the system is dimensioned so as to operate within astandard sized treatment room. In another aspect, the treatment patternis adapted so as to be suitable for use in a non-medical application,such as cargo scanning or non-destructive scanning.

In another aspect, the invention relates to a photon collimationassembly that includes: one or more photon generating layer; and asubstantially planar collimator block having an upstream side towards anelectron source when included in a treatment system for treating atargeted tissue and downstream side towards the targeted tissue, theupstream side being disposed adjacent the photon generating layer,wherein the collimator block includes a plurality of channels, eachextending from an inlet opening at the upstream side to an outletopening at the downstream side of the collimator block, wherein thechannels and outlet openings and a thickness of the block aredimensioned so as to suitably maintain resolution and contrast of anintensity-modulation pattern of a beam when collimated through thechannels. In some embodiments, suitably maintaining resolution of thepattern entails maintaining a resolution of the treatment pattern at theoriginal size within 1/10 of a width of the overall pattern or smaller,such as 1/100 of a width of the overall pattern.

In some embodiments, each of the channels of the collimation assemblyhas a substantially square cross-section throughout, while in otherembodiments, each of the channels has a non-square cross-sectionoptimized to produce specific beamlet shapes. The channels may bearranged on a rectangular or non-rectangular grid, the channels traversethe block at angles substantially perpendicular to the upstream anddownstream faces of the block, or the channels traverse the block atangles substantially oblique to the upstream and downstream faces of theblock. In one aspect, the assembly collimates without requiring movementof any mechanically moving parts.

In certain embodiments, a spacing between outlet openings of thechannels of the collimation assembly is sufficiently small that apenumbra of individual beams transmitted through the channels fills adosimetric gap in the targeted tissue between beamlets when used totreat the targeted tissue. In some embodiments, the spacing may be suchthat the adjacent beamlets overlap at the target. In one aspect, theopenings and thickness of the collimator are dimensioned so that apenumbra of individual beams transmitted through the channels issufficiently sharp to provide sufficient resolution to maintain anintensity-modulation pattern of the beams when transmitted through thechannels. In some embodiments, the outlet opening is substantiallylarger than the inlet opening for each of the plurality of channels.

In some embodiments, the collimation assembly includes a photongenerating target consists of individual target material plugs alignedover the corresponding channels of the collimator array embedded withina layer of heat conducting material situated at the upstream side of thecollimator array. In certain embodiments, the collimation assemblyincludes an active cooling feature, such as cooling through flow ofwater or air through a portion of the collimator or component thermallycoupled with the collimator assembly.

In another aspect, collimation assembly may be provided in a treatmentpositioned in beamline or may be include in a set of differingcollimation assemblies. In some embodiments, a treatment system includesa rotating gantry on which one or more beamlines are mounted along withone or more collimation assemblies such that collimated beams can bedirected to the target tissue from multiple directions by rotating thegantry. In another aspect, a system may include a carousel of differingcollimation assembly that rotates so as to position as selectcollimation assembly in a desired beamline for treatment.

In one aspect, the system includes a combination of steering magnets,permanent magnets or electromagnets and accelerator assembled in acompact instrument delivering medium range energy electron bunches witha controlled intensity profile at a desired central target, andduplicable to cover a minimum number of incoming angles (e.g. 16)distributed around the target, and contained within a standard treatmentroom size. Typically, a standard treatment room size, such as about20×20 feet wide and 10 feet high. Transport, magnification and deliveryof the desired electron beam at the central target is achieved bycombining steering, optical magnification, emittance preservation bymeans of a minimum number of magnets. In some embodiments, steering isperformed concurrent with magnification, thereby allowing furtherreduction in size of the system.

Delivery of radiation therapies in significantly reduced time-scale ascompared to convention methods poses a number of difficulties, many ofwhich are addressed by the methods and systems described herein. Forexample, aspects relating to targeted tissue motion, radiation beamgeneration and steering, power production and distribution, radiationsource design, radiation beam control and shaping/intensity-modulation,treatment planning, imaging and dose verification present variouschallenges and, as used in conventional therapies, barriers todelivering radiation therapies to targeted tissues on a significantlyreduced time scale. While the methods and systems described herein maybe used to facilitate very rapid radiation therapies, particularly byaddressing the above noted aspects of radiation delivery therapies, itis understood that these methods and systems are not limited to anyparticular radiation therapy delivery system or application describedherein, and may be advantageous when used in various other radiationtherapies and delivery systems, including conventional radiationtherapies as well as non-medical applications.

A fundamentally different approach to managing motion is to deliver thetreatment so rapidly that no significant physiologic motion occursbetween verification imaging and completion of treatment. According tocertain embodiments of the invention, an accelerator, more preferably acompact high-gradient, very high energy electron (VHEE) linearaccelerator, which may be a standing wave linear accelerator, togetherwith a delivery system capable of treating patients from multiple beamdirections, potentially using all-electromagnetic or radiofrequencydeflection steering is provided, that can deliver an entire dose orfraction of high-dose (e.g., 20-30 Gy) radiation therapy sufficientlyfast to freeze physiologic motion, yet with a better degree of doseconformity or sculpting than conventional photon therapy. The term“sufficiently fast to freeze physiologic motion” in this document meanspreferably faster than one human breath hold, more preferably less than10 seconds, even more preferably less than 5 seconds, even morepreferably less than one heartbeat and most preferably less than asecond. In addition to the unique physical advantages of extremely rapidradiation delivery, there may also be radiobiological advantages interms of greater tumor control efficacy for the same physical radiationdose. Certain embodiments of the invention can also treat non-tumortargets, such as, by way of nonlimiting example, ablation or othertreatment of: (1) nerves or facet joints for pain control; (2) foci inthe brain for neuromodulation of neurologic conditions including pain,severe depression, and seizures; (3) portions of the lung with severeemphysema; and/or (4) abnormal conductive pathways in the heart tocontrol refractory arrhythmias.

According to certain embodiments of the invention, there is provided asystem for delivering very high electron energy beam to a target in apatient, comprising: an accelerator capable of generating a very highelectron energy beam; a beam steering device capable of receiving thebeam from the accelerator and steering the beam to the target frommultiple directions; and a controller capable of controlling length oftime that the beam irradiates the target, the length of timesufficiently fast to freeze physiologic motion, and to control thedirections in which the beam steering device steers the beam to thetarget.

In certain embodiments, the controller is configured to receiveinformation from an imaging device and use the information from theimaging device to control the directions in which the beam steeringdevice steers the beam to the target. In some embodiments, theaccelerator is a linear electron accelerator capable of generating abeam having energy of between 1 and 250 MeV, more preferably 50 and 250MeV and most preferably between 75 and 100 MeV. In a rapid radiationtreatment embodiment, the time period is preferably faster than onehuman breath hold, more preferably less than 10 seconds, even morepreferably less than 5 seconds, even more preferably less than oneheartbeat and most preferably less than a second. The beam steeringdevice may include an electro-magnetic device and/or a radiofrequencydeflector device. In some embodiments, the beam steering device includesa gantry, the gantry including multiple beam ports. The beam ports maybe disposed in various arrangements, including arrangements that areannular, staggered, and planar, non-planar. In some embodiments, thebeam steering device includes a continuous annular gantry. In certainembodiments, the beam steering device is capable of providing thinpencil beam raster scanning.

Methods of utilizing beams of spatially varying beam intensity and thecollimation assembly features described above to provide a radiationtreatment, particularly a rapid radiation treatment, are also providedherein.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic representation of a system in accordance withcertain embodiments of the invention, showing beam access from a largenumber of axial directions by electromagnetic- or radiofrequencydeflection steering.

FIGS. 2A-2F show comparative simulation results of SABR for an earlystage lung tumor using 6 MV photons, 20 MeV electrons, and 100 MeVelectrons.

FIGS. 3A-3E show a schematic (FIG. 3A) and photograph (FIG. 3B) of theexperimental setup for film measurements (FIG. 3C) of very high energyelectron beams at the Next Linear Collider Test Accelerator (NLCTA) beamline at the SLAC National Accelerator Laboratory (SLAC), together withMonte Carlo simulations (solid lines) and film measurements (markers) ofpercentage depth dose curves (FIG. 3D) and beam profiles taken at 6 mmdepth (FIG. 3E) for 50 MeV and 70 MeV beams, respectively.

FIG. 4 shows graphic representations of percentage depth doses for a 2×2cm 100 MeV electron beam in a water phantom, simulated using threeindependent Monte Carlo codes.

FIG. 5 shows graphic representations of percentage depth doses for 2×2cm 50 and 150 MeV electron beams compared to 6 MV photons in a waterphantom, with 2 cm thick heterogeneous tissue at 10 cm depth.

FIG. 6 shows graphic representations of relative contribution to dosefrom a 100 MeV electron beam vs. secondary generated particles(logarithmic scale).

FIG. 7 shows water phantoms used in Monte Carlo simulations conducted inaccordance with certain embodiments of the invention.

FIGS. 8A-8E schematically show portions of a radiation treatment systemwith modulation of electron beam transverse profile using pulse-to-pulsemodulation of injection laser beam profile impacting a photocathode ofan electron injector.

FIGS. 9A-9B illustrate electric and magnetic fields in an examplephoto-gun in accordance with certain embodiments of the invention.

FIG. 10 illustrates associated electric fields in an example photo-gunin accordance with certain embodiments of the invention.

FIG. 11 illustrates an example photo-gun in accordance with embodimentsof the invention.

FIG. 12 illustrates a schematic of a beamline in an example treatmentsystem in accordance with certain embodiments of the invention.

FIGS. 13A-13B illustrate treatment patterns at the source and at thetargeted tissue, respectively, after steering and magnification of theintensity-modulated beam in accordance with certain embodiments of theinvention.

FIGS. 14A-14B illustrate examples of collimator assemblies for use in arapid radiation treatment system.

FIG. 15 illustrates a cross section of the channels through a collimatorassembly in accordance with certain aspects of the invention.

FIG. 16 shows photon fluence when one channel is illuminated by anelectron beam of 10 MeV energy in accordance with certain aspects of theinvention.

FIG. 17 shows the corresponding dose distribution at 1.5 cm depth in awater phantom in accordance with certain aspects of the invention.

FIG. 18A shows a conceptual schematic channel configuration in which thephoton beamlets diverge from one another. They are dosimetricallymatched without a gap in the plane of the targeted tissue in thisexample. Channel spacing may also be chosen such that the beamletsoverlap in the plane of the targeted tissue.

FIG. 18B shows a conceptual schematic channel configuration in whichclusters (pairs in this example) of channels converge to a focus in theplane of the targeted tissue, and different clusters form beamletclusters that diverge from each other. Other sophisticated geometriesare possible.

FIG. 18C shows a conceptual schematic channel configuration in whichchannels of various sizes are interspersed to provide for example largerarea coverage with larger channels and finer field edge shaping withsmaller channels.

FIG. 18D shows a conceptual schematic channel configuration of anexample collimator assembly in which the channels are largely at anoblique angle to the upstream and downstream faces of the block.

FIG. 18E shows a schematic of a treatment system utilizing the examplecollimator assembly in FIG. 18D.

FIG. 18F shows a schematic of a treatment system having a rotatinggantry that includes collimation assemblies, in accordance with aspectsof the invention.

FIG. 18G shows a treatment system having multiple beamlines from asingle accelerator for use with or without a collimation assembly, inaccordance with aspects of the invention.

FIG. 19A shows a bremsstrahlung target array comprising tungsten plugsaligned with the collimator channels, embedded in a layer of copper forheat dissipation and conduction.

FIG. 19B shows bremsstrahlung targets comprising multiple thin layers oftungsten distributed along the length of each channel with fixed orvariable spacing between layers.

DETAILED DESCRIPTION OF THE INVENTION

I. Rapid Radiation Treatment

A. Significance

In the U.S., cancer has surpassed heart disease as the leading cause ofdeath in adults under age 85, and of the 1.5 million patients diagnosedwith cancer each year, about two thirds will benefit from radiationtherapy (RT) at some point in their treatment, with nearly threequarters of those receiving RT with curative intent. Worldwide, theglobal burden of cancer is increasing dramatically owing to the agingdemographic, with an incidence of nearly 13 million per year and aprojected 60% increase over the next 20 years, and the number ofpatients who could benefit from RT far exceeds its availability.Moreover, even when RT is administered with curative intent, tumorrecurrence within the local radiation field is a major component oftreatment failure for many common cancers. Thus, improvements in theefficacy of and access to RT have tremendous potential to saveinnumerable lives.

Although there have been major technological advances in radiationtherapy in recent years, a fundamental remaining barrier to precise,accurate, highly conformal radiation therapy is patient, target, andorgan motion from many sources including musculoskeletal, breathing,cardiac, organ filling, peristalsis, etc. that occurs during treatmentdelivery. Conventional radiation delivery times are long relative to thetime scale for physiologic motion, and in fact, more sophisticatedtechniques tend to prolong the delivery time, currently 15-90 minutesper fraction for state-of-the-art high-dose radiotherapy. The veryfastest available photon technique (arc delivery with flattening filterfree mode) requires a minimum of 2-5 min to deliver 25 Gy. Significantmotion can occur during these times.

Even for organs unaffected by respiratory motion, e.g., the prostate,the magnitude of intrafraction motion increases significantly withtreatment duration, with 10% and 30% of treatments having prostatedisplacements of >5 mm and >3 mm, respectively, by only 10 minuteselapsed time. As such, considerable effort has been devoted todeveloping “motion management” strategies in order to suppress, control,or compensate for motion. These include complex immobilization, fiducialmarker implantation, respiratory gating, and dynamic tumor tracking, andin all cases still require expansion of the target volume to avoidmissing or undertreating the tumor owing to residual motion, at the costof increased normal tissue irradiation.

Several factors contribute to long delivery times in existing photontherapy systems. First, production of x-rays by Bremsstrahlung isinefficient, with less than 1% of the energy of the original electronbeam being converted to useful radiation. Second, collimation, andparticularly intensity modulation by collimation, is similarlyinefficient as the large majority of the beam energy is blocked bycollimation. Third, using multiple beam angles or arcs to achieveconformal dose distributions requires mechanical gantry motion, which isslow. Treatment using protons or other heavier ions has dosimetricadvantages over photon therapy, and these particles can beelectromagnetically scanned very rapidly across a given treatment field.However changing beam directions still requires mechanical rotation ofthe massive gantry, which is much larger and slower than for photonsystems. The cost and size of these systems also greatly limits theiraccessibility.

Very high-energy electrons (VHEE) in the energy range of 50-250 MeV haveshown favorable dose deposition properties intermediate betweenmegavoltage (MV) photons and high-energy protons. Without the need forinefficient Bremsstrahlung conversion or physical collimation, and witha smaller steering radius than heavier charged particles, treatment canbe multiple orders of magnitude faster than any existing technology in aform factor comparable to conventional medical linacs. According tocertain embodiments of the invention, a compact high-gradient VHEEaccelerator and delivery system is provided that is capable of treatingpatients from multiple beam directions with great speed, usingelectro-magnetic, radiofrequency deflection or other beam steeringdevices. Such embodiments may deliver an entire dose or fraction ofhigh-dose radiation therapy sufficiently fast to freeze physiologicmotion, yet with a better degree of dose conformity or sculpting, anddecreased integral dose and consequently decreased risk of latetoxicities and secondary malignancies, than the best MV photon therapy.Suitable energy ranges in accordance with certain embodiments of theinvention are 1-250 MeV, more preferably 50-250 MeV, and most preferably75-100 MeV. Again, as described in the Summary section above, the term“sufficiently fast to freeze physiologic motion” in this document meanspreferably faster than one human breath hold, more preferably less than10 seconds, even more preferably less than 5 seconds, even morepreferably less than one heartbeat and most preferably less than asecond.

According to some embodiments, a major technological advance isextremely rapid or near instantaneous delivery of high dose radiotherapythat can eliminate the impact of target motion during RT, affordingimproved accuracy and dose conformity and potentially radiobiologicaleffectiveness that will lead to improved clinical outcomes. Rapidimaging and treatment can also lead to greater clinical efficiency andpatient throughput. For standard treatments, the room occupancy time canbe reduced to less than 5 minutes. There can also be a great practicaladvantage for special populations like pediatric patients who normallyrequire general anesthesia for adequate immobilization during longtreatments, and who can instead be treated with only moderate sedationfor such rapid treatments. Such advantages can be achieved, according tosome embodiments, in a compact physical form factor and low costcomparable to conventional photon therapy systems, and much lower thanhadron therapy systems. One embodiment is shown in FIG. 1, which shows asystem wherein beam access from a large number of axial directions isachieved by electromagnetic steering without moving parts or with aminimum of moving parts, for extremely fast highly conformalradiotherapy. The system shown in FIG. 1 includes a compact linearaccelerator, a beam steering device, and a controller for controllingthe very high electron energy beam that is delivered to the patient. Theembodiment can also include an integrated imaging device that obtainsimages of portions of the patient including the tumor or other site tobe treated. The imaging device can also provide information to allow forcontrol of the beam steering device in order to control directions fromwhich the beam is delivered, and timing of the beam, among othervariables.

Furthermore, the prolonged treatment times of conventional highlyconformal RT are sufficiently long for repair of sublethal chromosomaldamage to occur during treatment, potentially reducing the tumoricidaleffect of the radiation dose. Thus in addition to the unique physicaladvantages of extremely rapid radiation delivery, there may also be doseadvantages. It is hypothesized that the treatment times sufficientlyfast to freeze physiologic motion that are made possible by certainembodiments of the invention may be more biologically effective,producing enhanced tumor cell killing for the same physical dose.Differences between certain embodiments of the invention andconventional photon therapy that impact biological effectiveness includea much faster delivery time and differences in the radiation quality.

Dose rate effects are well described in the radiobiology literature, inwhich prolongation of delivery times results in decreased cell killing.The main mechanism known to be responsible for this effect is repair ofpotentially lethal DNA double strand breaks (DSB) during the intervalover which a given dose of radiation is delivered. Several in vitrostudies have demonstrated significantly decreased cell killing whendelivery is protracted from a few minutes to tens of minutes. However,there is a lack of consensus in the literature regarding the kinetics ofsublethal damage (SLD) repair, with some studies suggesting thatcomponents of SLD repair may have repair half-times of as little as afew minutes. If so, shortening the delivery times even from a fewminutes to a time period sufficiently fast to freeze physiologic motionhas the potential to increase tumor cell killing.

B. Beam Steering

Some embodiments of the invention take advantage of the fact thatelectrons are relatively easier to manipulate using electric andmagnetic fields. Charged particles such as electrons and protons can beproduced as spatially coherent beams that can be steeredelectromagnetically or with radiofrequency deflection with highrapidity. Thus, direct treatment with scanned charged particle beams caneliminate the inefficiencies of Bremsstrahlung photon multiple beamsfrom different directions toward the target in the patient. Allconventional radiation therapy systems accomplish multidirectionaltreatment by mechanically rotating a gantry, or an entire compact linac,or even cyclotron, directing radiation to the target from one directionat a time.

As a preliminary matter, at the end of the accelerator structure thebeam must be deflected and then transported to the exit port and towarda target in or on the patient, such as a tumor in the patient. At theexit port the beam must be steered again to change the exit angle and/orbeam size to adapt to the treatment plan. Electro-magnetic and/or RFdeflector steering systems will manipulate the electron beam.

A variety of gantry designs are potentially available, from simple tocomplex, ranging from multiple discrete beam ports arranged around thepatient to a continuous annular gantry to allow arbitrary incident axialbeam angles. The design depends on a number of factors, includingscanning strategies such as thin pencil beam raster scanning vs. volumefilling with non-isocentric variable-size shots, or use of transversemodulation of the electron beam profile.

According to one embodiment, the steering system of the electron beamstarts at the end of the accelerator structure with a two-dimensionaldeflector, which guides the beam into one of multiple channels. Once thebeam enters a specific channel it is guided all the way to the exit ofthe channel, which is perpendicular to the axis of the patient. Theguidance through the channels is achieved using low aberration electronoptics. At the exit of each channel another small 2-D deflector can beadded to scan the beam over a target. The number of channels can then beabout 10-50. For a given channel width, a larger initial deflectionwould increase the number of channel entry ports that fit into thecircumference swept by the beam. Thus if the field strength wereincreased, the number of channels could be increased to 100 or more.

Because a linear accelerator will typically consume 50 to 100 MW of peakpower to achieve 100 MeV of acceleration, over a length of 2 to 1 mrespectively, potentially megawatt powered RF deflectors can beconsidered. These have the advantage of being ultra-fast and permitcapitalization on the RF infrastructure that is used for the mainaccelerator structure. In any event, the delivery system is preferablyoptimized to achieve high-dose treatment times sufficiently fast tofreeze physiologic motion.

Beam steering systems according to certain embodiments of the inventionadopt a design that uses a smaller number of discrete beam channels, forexample 3-10, that are mechanically rotated with the gantry around thepatient. The initial deflector at the exit of the accelerator rapidlysteers beams into the channels as they rotate. Although the ideal is toeliminate the need for any mechanical moving parts, some advantages ofthis design include: arbitrary rotational angular resolution despite afixed number of beam channels; reduced complexity and possibly costgiven the smaller number of beam channels needed to achieve equivalentangular coverage; and the larger space between beam channels which makesit more straightforward to incorporate an x-ray source and detectingarray for imaging, which when rotated provides integrated computedtomography imaging. The rate of mechanical rotation preferably providesfull angular coverage sufficiently fast to freeze physiologic motion.The greater the number of beam channels, the less rotational speedrequired to meet this condition as a general matter.

One innovation of certain embodiments of the invention is to eliminatemechanical gantry rotation, thus a beam steering system with nomechanical moving parts. One such embodiment is illustrated in FIG. 1,in which there is a gantry through which a charged particle beam iselectromagnetically steered or steered using radiofrequency deflectionto the target from any axial direction and a limited range ofnon-coplanar directions in addition. An alternative implementation is touse multiple discrete beam ports arranged radially around the patient,with the beam being steered through each of the ports to the target formultidirectional beam arrangements. Another alternative implementationis to have multiple accelerating structures, one for each of a set ofbeam ports arranged radially around the patient.

Such novel treatment system geometries and steering systems can greatlyenhance the treatment delivery speed of radiation therapy using any typeof charged particle. Combining it with high-energy electrons in the1-250 MeV range, more preferably the 50-250 MeV range, most preferablythe 75-100 MeV range, has the following additional advantages: (1)Conformal dose distributions to both superficial and deep targets inpatients superior to what can be achieved with conventional high-energyphoton therapy; (2) Compactness of the source and power supply, which byusing high-gradient accelerator designs such as those based wholly orpartially on accelerators developed or in development at the SLACNational Accelerator Laboratory (SLAC) as described further below canaccelerate electrons up to these energies in less than 2 meters; (3)Compactness of the gantry/beam ports compared to protons or ions becauseof the smaller electro-magnetic fields needed for electrons. Thisresults in a system of comparable cost and physical size to existingconventional photon radiotherapy treatment systems, yet with better dosedistributions and far faster dose delivery.

If treatment with photon beams is still desired, an alternativeembodiment is to incorporate in this geometry an array of high densitytargets and collimator grid in place of a single target/multi-leafcollimator combination, one per beam port in the case of discrete beamports, or mounted on a rapidly rotating closed ring and targeted by thescanned electron beam in the case of an annular beam port, in order toproduce rapidly scanned, multidirectional photon beams. While thisapproach may be subject to the inefficiency of Bremsstrahlungconversion, the speed limitations of conventional mechanical gantry andmulti-leaf collimator motions may be essentially eliminated. The mainpotential advantage of this implementation is that existing commercialelectron linacs in a lower energy range could be used as the source.

In addition to extremely rapid dose delivery, certain embodiments of theinvention naturally facilitate rapid image-guidance to ensure accuracy.By adjusting the energy of the scanned electron beam and directing it toan annular target or a fixed array of targets, with an appropriatelyarranged detector array, extremely fast x-ray computed tomography (CT)or digital tomosynthesis images can be obtained and compared topre-treatment planning images immediately before delivery of the dose.Alternative embodiments can include integration of more conventionalx-ray imaging or other imaging modalities, positron emission tomographyand other options described further below.

C. Monte Carlo Simulation Design Considerations

One approach in designing certain embodiments of the invention is toproceed using some or all of the following: (1) Monte Carlo simulationsto determine optimal operating parameters; (2) experimental measurementsof VHEE beams to validate and calibrate the Monte Carlo codes; (3)implementation factors for practical, cost-efficient and compact designsfor the systems; and (4) experimental characterization of keyradiobiological aspects and effects.

1. Monte Carlo (MC) Simulation

MC simulations of VHEE of various energies have been performed on asample case to estimate the range of electron energies needed to producea plan comparable to optimized photon therapy. Dose distributions werecalculated for a simulated lung tumor calculated on the CT data set ofan anthropomorphic phantom.

Specifically, an optimized 6 MV photon beam Volumetric Modulated ArcTherapy Stereotactic Ablative Body Radiotherapy (VMAT SABR) plancalculated in the Eclipse treatment planning system, and simplisticconformal electron arc plans with 360 beams using a commonly available20 MeV energy and a very high 100 MeV energy calculated with the EGSnrcMC code were compared. (See Walters B, Kawrakow I, and Rogers D W O,DOSXYZnrc, Users Manual, 2011, Ionizing Radiation Standards NationalResearch Council of Canada. p. 1-109, available online at(http://irs.inms.nrc.ca/software/beamnrc/documentation/pirs794/),incorporated herein by this reference).

FIG. 2A-F show axial images of simulation of SABR for an early stagelung tumor: dose distribution in an anthropomorphic phantom for astate-of-the-art 6 MV photon VMAT plan (FIG. 2A), a conformal electronarc plan using currently available 20 MeV electron beam (FIG. 2B), and aconformal electron arc plan using a 100 MeV electron beam as might bedelivered by an embodiment of the invention (FIG. 2C). A graphicalrepresentation shows dose volume histogram (“DVH”) of the planningtarget volume (“PTV”) (delineated in black in the axial images) andcritical organs: DVHs for 6 MV photons are shown in solid, 20 MeVelectrons in dotted, and 100 MeV electrons in crossed lines (FIG. 2D).The plans were normalized to produce the same volumetric coverage of thePTV by the prescription dose. While conventional 20 MeV electronsresults in poor conformity, the 100 MeV electron plan, even withoutoptimization, is slightly more conformal than the 6 MV photon VMAT plan.Simulating conformal electron arcs across an energy range of 50-250 MeV(FIG. 2E, 2F) demonstrates that both the high (100%) and intermediate(50%) dose conformity indices (CI100% and CI50%) as well as the meanlung dose and total body integral dose are superior for electronenergies of ˜80 MeV and higher for this selected clinical scenario. Withinverse optimization, superior plans with even lower electron energiesshould be possible.

As shown in FIGS. 2A-2F, the axial views of the dose distributionsdemonstrate that when all the plans are normalized to produce the samevolumetric coverage of the target, the dose conformity of the 20 MeVbeam is poor whereas the 100 MeV electron beam, even without inverseoptimization, generates a dose distribution equivalent to thestate-of-the-art 6 MV photon beam VMAT plan. In fact, the DVH's of thetarget and critical structures for the three beams demonstrate slightlybetter sparing of critical structures with the 100 MeV electron plancompared to the 6 MV photon plan. As shown in FIGS. 2E and 2F, atelectron energies above ˜80 MeV, simple conformal electron arc plans(normalized to produce the same volumetric coverage of the target) aresuperior to the optimized 6 MV photon VMAT plan in terms of conformity,with conformity index defined as the ratio of the given percent isodosevolume to the PTV, and the normal organ doses (mean lung dose) and totalbody integral dose (expressed in arbitrary units normalized to thephoton plan). In preliminary simulations of this selected clinicalscenario, the inventors have found electron energies of 75-100 MeV toproduce plans of comparably high to superior quality compared to thebest photon plans, and anticipate that plan optimization will producesuperior plans with even lower electron energies. For example, theinventors have used Monte Carlo simulations to demonstrate that an 8 cclung tumor could be treated with 100 MeV electrons to a dose of 10 Gy in1.3 seconds.

Further optimization of the electron plan can help to define the minimumelectron beam energy with a comparable dose distribution to the bestphoton VMAT plan. In preliminary simulations of this selected clinicalscenario, the inventors have found electron energies of 75-100 MeV toproduce plans of comparably high quality to the best photon plans, andanticipate superior plans with plan optimization.

2. Experimental Measurement of VHEE Beams

a. Monte Carlo Simulations

To demonstrate the accuracy of Monte Carlo calculations with VHEE beams,the inventors experimentally measured the dose distribution and depthdose profiles at the NLCTA facility at SLAC. Of note, the NLCTA employscompact high-gradient linear accelerator structures which can producebeams that are relevant to those potentially suitable for certainembodiments of the invention. The inventors assembled a dosimetryphantom by sandwiching GAFCHROMIC EBT2 films (International SpecialtyProducts, Wayne, N.J.) between slabs of tissue equivalent polystyrene asshown in FIGS. 3A-3E. FIG. 3A is a schematic and FIG. 3B is a photographof the experimental setup for film measurements (FIG. 3C) of veryhigh-energy electron beams at the NLCTA beam line at SLAC. Monte Carlosimulations and film measurements of percentage depth dose curves (FIG.3D) and 2-D dose distributions taken at 6 mm depth (FIG. 3E) for 50 MeVand 70 MeV beams demonstrate a high degree of agreement betweencalculation and measurement.

By way of procedure and in greater detail, the phantom as shown in FIG.3A was irradiated with 50 MeV and 70 MeV beams. Three beam sizes rangingfrom 3.35 to 6.15 mm were tested for each energy level. The energy wasmeasured by a spectrometer upstream from the location of the experimentand the beam size was measured by two scintillating screens using twocameras just before and after the phantom with the phantom removed fromthe beam line (FIG. 3B). The films were calibrated with a clinicalelectron beam at 12 MeV. MC simulations have demonstrated no energydependence of the film response at electron energies above 1 MeV. Thenumber of particles required to irradiate the films to dose levelsbetween 1-5 Gy to match the dynamic range of the film was determined foreach beam size using MC simulations and used in the experiment. Thecharge was set to 30 pC/pulse corresponding to 1.9×10⁸ electrons and thepulse rate was reduced to 1 Hz for easier control of the exposure. Thenumber of pulses varied from 2 to 40 pulses depending on the beam size.The experimental and calibration films were read out in a flatbedscanner (Epson Perfection V500, Long Beach, Calif.) with 0.1 mm pixels24 hours after irradiation (FIG. 3C) and central axis percentage depthdose (PDD) curves and 2-dimensional dose distributions at various depthswere plotted. The experimental setup was simulated in MCNPX 5.0 MC code.(See Palowitz D B, MCNPX User's Manual, Version 2.7.0, 2011. availableonline at (http://mcnpx.lanl.gov/documents.html), incorporated herein byreference).

The simulations are compared to measurements in FIG. 3D-3E. Goodagreement was observed for both the PDD curves and beam profiles for 50and 70 MeV. These preliminary results indicate that dose from VHEE beamscan be measured with GAFCHROMIC films and that VHEE beams can beaccurately simulated with the GEANT4 code.

In the arrangement shown in FIG. 3B, a 50-μm vacuum window made ofstainless steel was used to interface the accelerator line with openair, in which the dose phantom (FIG. 2A) was placed. The stainlesswindow was found to cause significant angular beam spreading, so thatthe simulations were also performed with a beryllium window whichimparted less beam spreading. While a vacuum window is necessary toseparate the vacuum of the accelerator beam line from the open air andthe patient, significant angular spread will adversely affect beamperformance and clinical accuracy. The angular spread from a thinnerberyllium window was still present but it was much smaller than steel,due to beryllium's low atomic number.

b. Cross Validation of Monte Carlo Codes

The inventors performed Monte Carlo simulations using three independentcodes for identical geometries to determine the consistency ofcalculated doses. The dose deposition of a number of rectangularelectron beams incident on a 20×20×30 cm water phantom (as shown in FIG.7) was simulated in the GEANT4, MCNPX, and EGSnrc MC codes. Thesimulated electron beam energies were 50, 75, 100, and 150 MeV with beamsizes of 1×1 cm and 2×2 cm. The central-axis PDDs were plotted andcompared for all three MC codes. Excellent agreement was found betweenthe codes for all of these comparisons, as shown in FIG. 4, which showsPDD for a 2×2 cm 100 MeV electron beam, simulated using the three MonteCarlo codes.

c. VHEE Tissue Interactions

Monte Carlo simulations were performed to evaluate the impact of varioustissue heterogeneities on VHEE beams relative to MV photon beams. FIG. 5shows PDD curves for 2×2 cm 50 and 150 MeV electron beams compared to 6MV photons in a water phantom with 2 cm thick heterogeneous tissue at 10cm depth, normalized to identical dose at 3 cm depth. As shown in FIG.5, the 50 and 150 MeV VHEE beams are less sensitive to tissueheterogeneity over the density range from lung tissue to titaniumprosthetic implants compared to 6 MV photons.

Contribution of secondary particles produced by Bremsstrahlung andelectronuclear interactions to the dose from VHEE beams were alsoanalyzed. FIG. 6 shows relative contribution to dose from a 100 MeVelectron beam vs. secondary generated particles (log scale). As shown inFIG. 6, for a 100 MeV electron beam, nearly all the deposited dose isdue to electrons, with a minor contribution from Bremsstrahlung x-rays,and far lower dose from protons and neutrons. FIG. 6 also shows thatdose from neutrons is far less than with 15-18 MV photons or high-energyprotons. This holds for 50 and 70 MeV electrons as well (not shown). Fora 25 Gy SABR treatment of a 2 cm diameter target, an upper limit oftotal body neutron dose is estimated to be 0.6 mSv based on MCsimulations. This is in contrast to more than 1-2 orders of magnitudegreater estimated neutron doses of 9-170 mSv for scanning beam protontherapy and 15-18 MV photon IMRT for the same clinical scenario, basedon published measurements of ambient neutron doses [Schneider U, AgosteoS, Pedroni E, and Besserer J., “Secondary neutron dose during protontherapy using spot scanning,” International Journal of RadiationOncology Biology Physics, 2002; 53(1): 244-251. (PMID: 12007965); HowellR M, Ferenci M S, Hertel N E, Fullerton G D, Fox T, and Davis L W,“Measurements of secondary neutron dose from 15 MV and 18 MV IMRT,”Radiation Protection Dosimetry, 2005; 115(1-4): 508-512. (PMID:16381776) both of which are incorporated herein by this reference]. Anadvantage of such potential designs according to certain embodimentscompared to >8 MV photon and scanning beam or passive scattering protontherapies is elimination of need for beam modifying structures prior tobeam incidence on the patient, in which most neutrons are generated withexisting modalities.

d. Tissue Inhomogeneities

The effect of tissue inhomogeneities on dose deposition of VHEE beamshas been studied by the inventors. A 20×20×25 cm3 water phantom with0.5×0.5×0.1 cm3 voxels and a 2-cm thick inhomogeneity placed at 10 cmdepth was built (FIG. 7). The 2-cm thick slab was consequently filledwith lung with mass density ρ of 0.368 g/cm3, adipose (ρ=0.950 g/cm3),ribs (ρ=1.410 g/cm3), and cortical bone (ρ=1.920 g/cm3) tissue to assessthe effect of human tissue inhomogeneities. The tissue composition wasobtained from the ICRU-44 document [ICRU. Tissue substitutes inradiation dosimetry and measurement, 1989 (incorporated herein by thisreference)]. Moreover, the effect of metals, such as hip prostheses,dental fillings, and surgical clips, was investigated by simulating asteel slab (ρ=8.030 g/cm3). Doses deposited by 50, 100, and 150 MeVelectron beams, as well as 6 MV photon beam interacting with theinhomegeneity slab were simulated. The DOSXYZnrc code was chosen forthis task due to its simplicity of use and its shortest calculationtimes. The statistical uncertainties in all central axis voxels werebelow 1%.

3. Ultra-High Gradient Accelerator Structure Design

Pluridirectional very high electron energy radiation therapy systems andprocesses according to various embodiments of the invention can becreated with various types of electron source. There are a number ofpotential sources of very high-energy electrons in the range of, forexample, up to about 250 MeV. A non-exhaustive list includes cyclotrons,synchrotrons, linacs (which can include more conventional designs withgreater length), racetrack microtrons, dielectric wall accelerators, andlaser plasma wakefield accelerator sources. Some of these are large andwould need to be housed in a separate room. Some are not very maturetechnologies. In terms of goals of certain embodiments of the inventionwhich can include any or all of compactness (entire system fittingwithin existing medical linac vaults without a separate room), powerrequirements, cost, repetition rates, compatibility with intensitymodulation techniques described in this document, and other practicalconsiderations, compact very high-gradient standing wave linearaccelerators such as those developed at SLAC as described in the twoparagraphs immediately below, or derivatives of them, may be at least alogical starting point, although other currently existing or futureoptions should not be ruled out.

Highly efficient it-mode standing wave accelerator structures have beendeveloped at SLAC for the project formerly known as the Next LinearCollider, a positron-electron collider at 500 GeV energy for high-energyphysics research [Dolgashev V, Tantawi S, Higashi Y, and Spataro B,“Geometric dependence of radio-frequency breakdown in normal conductingaccelerating structures,” Applied Physics Letters, 2010; 97(17).(http://apl.aip.org/resource/1/applab/v97/i17/p171501_s1) incorporatedherein by this reference (hereinafter sometimes “Dolgashev 2010”). Suchaccelerators are capable of accelerating electrons to 100 MeV within 1meter (Id.) using an optimized accelerating waveguide powered by a 50 MW11.4 GHz microwave generator (klystron) [Caryotakis G. Development ofX-band klystron technology at SLAC. Proceedings of the 1997 ParticleAccelerator Conference, 1997; 3: 2894-2898.(http://ieeexplore.ieee.org/xpls/abs_all.jsp?arnumber=752852)incorporated herein by reference. In order to produce a practical systemin terms of cost and size, optimized designs according to certainembodiments of the invention allow both economical production and highperformance to minimize the treatment time while allowing maximumpossible flexibility in beamlet shapes, directionality, and energy.

Furthermore, it has been shown that coupling a series of small sectionsof standing-wave accelerators with a distributed radiofrequency (RF)network makes it possible to design a system without any reflection tothe RF source [Tantawi S G, “rf distribution system for a set ofstanding-wave accelerator structures,” Physical Review SpecialTopics-Accelerators and Beams, 2006; 9(11)(http://prst-ab.aps.org/abstract/PRSTAB/v9/i11/e112001) incorporatedherein by this reference (hereinafter, “Tantawi 2006”). Building onthese developments, practical implementations of a standing-waveaccelerator structure have been designed to accelerate electrons to 100MeV within one meter. (See for example, Neilson J, Tantawi S, andDolgashev V, “Design of RF feed system and cavities for standing-waveaccelerator structure,” Nuclear Instruments and Methods in PhysicsResearch A: Accelerators, Spectrometers, Detectors and AssociatedEquipment, 2011; 657(1): 52-54. (hereinafter, “Neilson 2011”), availableonline at(http://www.sciencedirect.com/science/article/pii/S0168900211008898),incorporated herein by reference). Such accelerators can serve as abasis for or be relevant to certain embodiments of the invention.

D. Other Design Issues

1. Design Options for the Injector System

To inject the required low charge bunch into accelerators according tocertain embodiments of the invention, several possibilities areavailable. Those include a photo-injector RF gun. Additional options canbe considered to reduce the cost and size of the system, including avariety of field emitter configurations and RF thermionic guns and DCphotocathode guns.

2. Optimization of the RF Source by the Addition of a Pulse CompressionSystem

RF source requirements depend ultimately, at least in part, on theaccelerator design. With the optimized cavities as described above, itis projected that a 50 MW source at X-band will be sufficient for a 2meter accelerator operating at 50 MV/m. This type of source is availableat SLAC and is being commercialized by Communications & Power Industries(Palo Alto, Calif.). With the use of a pulse compression system it maybe possible to either reduce the cost and sophistication of the RFsource dramatically or make the accelerator structure more compact byreducing the length to 1 meter. Because the typical filling time of sucha structure is about 100 ns and the RF source typically provides severalμs long pulses, one can use a compact RF pulse compressor with a highcompression ratio and a power gain of about 3.5 to reduce the requiredRF source power to only about 14 MW, which opens the door for a varietyof sources, including sources that are commercially available now, andincluding those that include a pulse compression system.

3. Imaging and Target Position Verification Options

Given that treatment according to certain embodiments of the inventionis delivered sufficiently fast to freeze physiologic motion, it isimportant to verify that the target is in the planned position at thetime the treatment is triggered or administered. Several dynamic or“real-time” imaging or other localization technologies can be integratedinto certain embodiments of the invention for this purpose. Potentialsuch implementations can include any of the following, alone or incombination: integration of two or more x-ray fluoroscopic imagingdevices; dynamic optical surface scanning; integration of fast x-raycomputed tomography; implantable radiofrequency beacons, whose3-dimensional position can be read out in real time by an externalantenna array. Beacons can be implanted in or near the target and serveas surrogates for the target position; MRI imaging and any of theapproaches described in PCT Application No. PCT/US2014/055270 filed Sep.11, 2014.

4. Implementation of Intensity Modulation

According to certain embodiments of the invention, which may be usedwith various types of accelerators in accordance with the invention, andin order to achieve highly conformal volumetric dose shaping, radiationfields from each of multiple beam directions can cover an area withvarying beam intensity across the field, with the intensity patternsoptimized to produce the desired 3-dimensional dose distribution whensummed across all beam directions. Such intensity modulation may beproduced by raster scanning individual beamlets of varying intensityacross the field from each beam direction. Alternatively, it may beproduced by using a 2-dimensional intensity-modulated electron patternat the source, effectively a simultaneously generated array of beamletsof varying intensity, and accelerate and steer the entire array to thetarget volume. This eliminates the need for a raster scanning mechanismat the exit of each of the beam channels, greatly simplifying the designand reducing the bulk and cost of those components, and increases thetreatment delivery speed by delivering beamlets in parallel within amuch smaller number of electron pulses or bunches.

II. Technologies to Facilitate Radiation Delivery in Rapid RadiationTreatments

A. Photo Cathode/Photo Electron-Gun

In accordance with certain aspects, methods and systems for rapidgeneration and delivery of transversely patterned electron beam totargeted tissue for rapid radiation treatment utilize a photo-electrongun. A photo-electron gun is one of various possible techniques that maybe used for precise and ultrafast dose delivery using a medical electronaccelerator in accordance with the present invention. The dose isproduced in rapid pulses of electrons delivered to the targeted tissuefrom different directions, different transverse beam pattern in eachdirection. Each pulse has a pre-programmed transverse dose pattern suchthat the total 3D dose conforms to the target volume in the patient.Projecting a pre-programmed light pattern on a photocathode generatesreplica of this light pattern with similar transverse distribution ofthe electrons. This pattern or image is then accelerated through lowaberration electron optics toward the targeted tissue.

In an exemplary embodiment, the system operates as follows: a lightsource, e.g laser, lamp, diode generates short, transversely uniformlight pulses. This light pulse is transversely modulated to producepre-programmed pattern using computer-controlled light-modulator. Thispatterned light pulse impacts photo-cathodes of the electron gun,creating electron-bunch replica of the light pattern. The resultingpatterned electron bunch is accelerated, and projected with requiredmagnification to create the desired intensity profile on the targetedtissue. In one aspect, any distortion of the pattern after steeringand/or magnification must be within acceptable limits so that thedesired intensity profile is delivered to the targeted tissue accordingto the treatment pattern, typically with spatial positioning accuracy ofwithin 3 mm and intensity within 3% of specified dose at the targetedtissue.

In accordance with various embodiments, the setup includes followingparts: light generation, its transport and diagnostics; photocathodeelectron gun; low aberration electron optics and electron beamdiagnostics.

According to some embodiments, the intensity modulation of the electronsource may be produced by using a photocathode illuminated by a lightsource with the corresponding intensity pattern, in effect, an opticalimage. One implementation is to use a laser as the light source, and adigital light processing (DLP) micromirror array or other intensitymodulating device to produce the charge image on the photocathode to beaccelerated and steered. The electron beam optics can be designed tomaintain the pattern with high fidelity until it reaches the target.

According to one approach, shown in FIG. 8A, a short, typicallypicosecond-long pulse with substantially uniform transverse profile isgenerated by a laser 1 (or at least a transverse profile having anintensity distribution that is appropriate to produce the desiredpattern at the targeted tissue according to the method described below)The wavelength of the laser is matched with specific photocathodematerial to obtain required charge and emittance. The laser pulse 2reflects off a digital-micro-mirror device 3. Pixels of thismicro-mirror device are controlled by a computer and will reflect aportion of the laser pulse 4 thus creating an image that is thentransferred to the photocathode 6 using precision projection optics 5.Although various types of accelerators may be used with this embodiment,high gradient pulsed devices with a few milliseconds between pulses arepreferable. The computer modulates the mirror array thus creating a newimage for each consequent pulse. A laser pulse with amplitude-modulatedtransverse profile that impacts the photocathode 6 will create anelectron replica of the laser pulse transverse profile 8. Thephotocathode 6 is a part of photo-electron gun 7. The gun creates anelectric field on the photocathode which accelerates thetransverse-modulated electron beam. The gun also provides initialacceleration to boost the electrons to relativistic velocities. Theelectron beam then passes through the low-aberration focusing systemtoward accelerator 10. The accelerator increases energy of the beam to adesired value. The electron beam then passes through focusing optics 11toward horizontal 12 and vertical 13 deflectors. The deflectors arecontrolled by a computer and are able to send the electron beam indifferent directions for each consecutive accelerator pulse. The desireddirection will depend on (among other things) specific realization ofthe gantry's beam lines, number of the beam lines and whether they aremovable or not. For clarity only one gantry beam line is shown in FIG.8A. After the deflectors, the electron beam passes through bendingmagnets 14, 16, 18 and electron optics 15, 17 and is directed throughelectron-beam monitoring system 19 toward the target 20. Thetransversely modulated electron beam irradiates the target with requireddistribution of the dose. After passing through the target, the beam issent toward beam dump 21 in order to reduce unwanted radiation exposureof the target.

In certain other embodiments, fast deflectors may be used to scan apatterned beam so as to create an effectively larger pattern, whichallows the coverage to be increased such as coverage up to 40×40 cm. Incertain additional other embodiments, beam deflectors may be used todirect the beam from a single accelerator to more than one beamline,each providing a beam direction toward the targeted tissue. For example,a system having 16 beamlines may require only four electron guns andaccelerators if each accelerator feeds four beamlines.

In another approach, the beams having an intensity-modulation directedto the target tissue according to a treatment pattern may be generatedby rastering of the beam. FIG. 8B illustrates the concept of a radiationtreatment system with pulse-to-pulse modulation of electron beamintensity and rastering of the electron beam at low energy. In thisembodiment, a pulsed injection laser 101 generates pulses of laser beams102, all with the same intensity after which a sub-nanosecond laser beamintensity modulator 103 controlled by a computer forms laser beams withmodulated intensity 104 according to the same transverse profile, suchthat it may be generated by a single spot laser beam. The laser beamshaving modulated pulse-to-pulse intensity 104 impact the photocathode105, creating an electron replica of the laser pulse according to thetransverse profile. The photocathode 105 is part of a photo-electron gun106 which creates an electric field on the photocathode that accelerateselectron bunches with same transverse profile but with modulatedintensity 108 to the fast rastering transverse deflector with verticalkick 109 and the fast rastering transverse deflector with horizontalkick 110 producing low-energy electron bunches with modulated intensitythat are rastered in vertical and horizontal directions 111. Theselow-energy electron bunches are then accelerated in a high gradientelectron accelerator 112. The accelerated electron bunches 113 passthrough a low aberration lens 114 and horizontal plane and verticalplane magnetic deflectors 115, 116 and are then steered with dipolemagnets 117, 119, 121 and focused with low aberration electron lenses118, 120 for delivery to the targeted tissue. An electron beam monitor122 may be used to monitor the beam to verify the intensity-modulationaccording to the treatment pattern before delivery to the target tissue123. Any remaining beam passing through the tissue is absorbed byelectron beam dump 124.

FIGS. 8C through 8E depict various aspects of the rastering approach inFIG. 8B. FIG. 8C illustrates a modulated charge emitted fromphoto-cathode of photo-electron gun after it exposed to laser pulseswith modulated intensity (item 108 in FIG. 8B); FIG. 8D illustrates aplot of vertical and horizontal momenta of electron bunches afterpassing through two fast rastering transverse deflectors, one withchanging vertical kick, another one with changing horizontal kick (item111 in FIG. 8B). Note that this is but one rastering solutions out ofmany, as one of skill in the art would be aware. FIG. 8D illustrates aplot of vertical and horizontal coordinate of electron bunches on atargeted tissue (item 123 in FIG. 8B)

Using the same accelerator and beamline transport and magnificationoptics as required to transport a spatially patterned electron source tothe targeted tissue, the same effect can be achieved by modulatingcharge and transversely rastering an electron point source at lowenergy, prior to the high energy accelerator. As examples, the electronpoint source can be created by producing photo-electron point sources ata fixed location on a photocathode, with modulated laser intensitychanging pulse-to-pulse charge according to a pre-programmed function,then using fast deflectors to raster the pulses. Another method ofcreating a rastered electron source is scanning the laser spot withmodulated intensity across the photo-cathode to produce requiredtransverse pattern but over multiple pulses. The difference from theconcept described in section I.C.3: Ultra-high gradient acceleratorstructure design is that the spatial modulation of dose in the targetedtissue will be produced by sequential irradiation by multiple pencilbeams with pre-programmed intensity, as opposed to irradiating anextended field simultaneously with single or many patterned beams.

Of note, a greater degree of intensity modulation will produce moreconformal dose distributions. However, with conventional photon therapywhere intensity modulation is delivered in a serial fashion over time,more modulation comes at a cost of longer delivery time, more leakagedose to the patient, and greater uncertainty in delivered dose becauseof target and organ motion during the longer treatment delivery time andits interplay. With VHEE technology according to certain embodiments ofthe invention, all of these problems are circumvented: arbitrarilycomplex intensity modulation can be produced through optical imaging,and rapid parallel delivery eliminates uncertainty from interplayeffects.

The concept of conversion of an optical intensity pattern into aradiation intensity pattern within a patient is considered to be unique,and also uniquely applicable to electron beam therapy in accordance withembodiments of the invention as opposed, for example, to photon orproton or other particle therapies. In general, the light-pulsegeneration could be based on laser, light-emitting diode, or variousother light sources with power, wavelength, and pulse length optimizedto produce sufficient electron charge and initial emittance from aspecific photocathode material.

In some embodiments, the pre-programmed light pattern is imprinted onthe light pulse using programmable transparent 2D screens, for exampleLCD, modulated using programmable mirror arrays (similar to ones used invideo-projectors), programmable micro-lens arrays, or programmabledeformable mirrors. Using different approach, the light pattern could begenerated directly in the light-source from programmable laser orlight-emitting-diode 2D array. These programmable devices changetransverse light pattern many times per second. Each pattern issynchronized with pulses of electron accelerators in order to generatetreatment dose with different transverse pattern on every consequentpulse, again, multiple times per second. The diagnostics of the lightpattern is performed simultaneously with the projection of the light onthe photocathode, by either splitting light pulse before projection onthe photocathode or by imaging the light pulse reflected from thephotocathode.

In accordance with the above embodiment, to extract photo-electrons fromthe photocathodes, the photo-gun exhibits high electric field on itssurface, which may be provided by any of DC high voltage, pulsed highvoltage, radiofrequency electromagnetic fields or any combinationthereof. The photocathode may be transparent or reflective, such as thephoto-gun 7 with metal reflective photo-cathode 6 shown in FIG. 10. Thematerials of the photocathode may include a metal (e.g. copper,magnesium, silver, or other suitable material), dielectric (e.g.diamond), semiconductor (GaAs, etc.), or various compound materialssuitable for use in photo-multipliers, streak cameras, various Cscompounds like CsBr, Sb—Cs or materials with a suitable energy gap. Thespecific photocathode material can be matched in quantum efficiency andlight-wavelength to the specific light source. For example UV lasercould be used with copper cathode, or CsBr cathode could be used withblue-light laser diode.

In accordance with the above embodiments, the patterned light pulseimpacting the photo-cathode creates an electron replica, which isaccelerated out of the photo-gun. The electro-magnetic fields of the guncan be configured to preserve phase space of the patterned electronbunch. Prior to delivery, the patterned bunch may be monitored along itspath to the targeted area, namely during acceleration, transport andconditioning, by means of invasive electron-bunch diagnostics likephosphor screens, wire scanners, and the like. By use of diagnostics,any drift in the tuning of the instrument to deliver the appropriatepattern to the targeted tissue can be corrected for and a high qualityof the delivery of the beam to the targeted tissue can be insured.Monitoring of the electron bunch properties during delivery may beperformed with non-invasive diagnostics such as beam position monitors,phase-cavities 10′, 10″ (shown as single cells at the entrance and exitof linac 10 in FIG. 8A), ion chambers and the like. The non-invasivediagnostics may record centroid transverse and temporal properties andcharge at the time of delivery.

In accordance with the above embodiments, after exiting the gun, thebunch is transported in low aberration electron optics and acceleratedin electron-accelerator to required energy 1-250 MeV. The transport andthe accelerator are designed to preserve emittance of the bunch by meansof low-aberration components. The property of the electron-beamtransport allows creation of the pre-programmed dose pattern on thetargeted tissue.

FIGS. 9A-9B illustrates electric and magnetic fields in an example 9.3GHz photo-gun suitable for use with the above described embodiments(only one-quarter of the photo-gun being shown). FIG. 9A illustratessurface electric fields and FIG. 9B illustrates surface magnetic fields.FIG. 10 illustrates the 9.3 GHz photo-gun and associated electric fieldsshown in FIG. 11 and further illustrates the various components anddirection of the patterned light pulse suitable for use with the abovedescribed embodiments. In some embodiments, the photocathode 6 on whichthe source treatment pattern is formed includes a CesiumBromide (CsBr)coating. In this example, the peak electric field on the cathode is 164MV/m for 2 MW of input power. The electric fields can be optimized topreserve bunch emittance. The advantages of using such a design fordelivering the radiation include rapid and precisely shaped dosedelivery far beyond that of conventional technologies. In addition, thisconfiguration allows for dose modulating and shaping without requiringmovement of mechanically moving parts, thereby allowing far more rapiddelivery than would be possible with delivery systems used inconventional radiation treatments. This allows delivery of a precise 3Ddistribution of whole-treatment dose in reduced treatments times, suchas delivery of an entire treatment dose in less than a minute, such asless than 10 seconds, preferably one second or less, or even morepreferably delivery in a sub-second time, which is not feasible withconventional treatment systems.

The above described embodiments utilizing a photo-gun are particularlysuited for treatments with high-energy electron beams, although it isunderstood that this feature may be used with various other radiationtreatment systems including treatment with X-rays. In one such example,a transverse modulated electron beam may be directed to a bremsstrahlungX-ray target combined with X-ray collimators. This combination allowsformation of a transversely patterned X-ray dose, which allows deliveryof an entire treatment dose in a very short time.

In one aspect, use of a photo-gun, as described in the aboveembodiments, provides improved rapid, pulse-to pulse modulation of thetransverse beam for use in medical accelerators. A medical acceleratorwith high-energy (50-250 MV) electron beam produces sufficient treatmentdose using relatively small charge well within the parameter ranges ofthe present invention. As one of skill in the art would understand, theabove described feature utilizing a photo-gun may be used in variousdiffering radiation systems, including those described herein. Thesource should be able to produce a spatially modulated pattern ofelectron emission, with low transverse emittance and longitudinalemittance (compatible with the required compactness of the beamlineoptics and with the RF frequency). Examples include RF powered guns withphotocathodes (currently the most practical option), but other designsfor patterned electron emission sources are possible and may become morepractical in the near future.

In certain aspects, the patterned electron beam is accelerated,transported, and magnified to the appropriate size to cover the targetedtissues in the patient. The shape, resolution, and contrast of thepattern must be preserved throughout this process. The size of thepattern on the source is typically a few millimeters and thus thesmallest feature size or resolution must be a few micrometers. To coverfield sizes treated in general radiation therapy (typically up to 30-40cm), the degree of beam magnification by the time it reaches the patientis on the order of 100-200, and may be different in the two transversedirections if desired.

FIG. 12 shows one possible compact beamline implementation (only onebeamline is shown for clarity) in which the entire system fits within atypical size radiation therapy vault. A series of magnets is included tobend and magnify the beam after acceleration through the linac 10. Inthe embodiment shown, the linac is 1-m long and accelerates theelectrons to form a 100 MeV electron beam, after which the beam passesthrough focusing optics 7, such as a series of quadrupole magnets, thenis steered by use of bending magnets 14, 16 alternating with electronoptics 15 to direct the electron beam to the targeted tissue in thepatient P. This beam dynamics simulation shows that an electron patternof less than 2 mm at the source with minimum feature size of about 10micrometers can be accelerated, transported (including bending towardthe patient), and magnified to a final size of nearly 20 cm in thisexample (about 100 times magnification), with preservation of thepattern including the finest features, such as shown in example patternof FIG. 13A). FIG. 12 shows an example of a single angle delivery systemat the most constraining bending angle (90 degree) from beam generation,acceleration, transport with steering and magnification; the last magnetends at 60 cm from the patient axis. While various embodiments may bedescribed as having certain dimensions, it is appreciated that theseconcepts may be used in systems having various other dimensions andconfigurations to obtain the advantages described herein.

In one aspect, the pattern imprinted on the photo-generated electronbeam at the emission from the cathode is chosen to match the requiredpattern needed at the patient targeted tissue, compensating foraberrations in the beam source, accelerating and transport systems. Thesteering system and magnifying system of the multiple beamlines occupiesa compact space (for example less 3 meters, the height of existingtherapy vaults). In some embodiments, a first very strong small apertureelectromagnetic lens acts on the small size beam exiting theaccelerator. The subsequent electromagnetic lenses have lower strength,compatible with larger apertures as the beam starts expanding. Thiscombination makes all electromagnetic lenses small enough to accommodatethe multiple beamlines. In some aspects, steering may be performed byone or more focusing elements, which may include permanent magnets,electro-magnets, electro-magnetic lenses or combinations thereof.Typically, dynamic steering of high energy beams is performed byelectro-magnetic deflection, that is forces generated by both electricand magnetic fields.

In certain embodiments, a beam spreader (12,13 in FIG. 8A and 115,116 inFIG. 8B) can be used after the accelerating structure to feed inmultiple distribution lines up to a few kHz rate to deliver beam to thepatient from the various angles using a single common accelerator, thussimplifying the RF distribution system and the footprint. For this case,the laser generation system rapidly (again few kHz rate) switches themodulation pattern sent to the single photocathode matching thedifferent patterns needed at the multiple delivery angles.

In some embodiments, a rapidly changing two-axis electromagnetic steerercan be added after the last electromagnetic lens to cover up to 40×40 cmof targeted tissue. Typically, the time needed to switch this steererstrength is less than the time elapsed between 2 consecutive deliveriesto the same arm of the multiple angle arms. Variations of this rasteringsolution can be used to relax the requirement on magnification to 50 orless. It is appreciated that the 40×40 cm² coverage noted above is butone example of a desired maximum field size that may be typical for aphoton treatment system and that various other configurations providingcoverage for larger areas may be realized depending on the application.For example, the technologies described herein may be utilized forvarious other applications, including but not limited to cargo scanningand non-destructive scanning (e.g. testing/analysis of materials orcomponents in bridges or buildings). Application of these technologiesfor such applications may necessitate variations in the examplesdescribed herein suited for the particular application of interest.

FIGS. 13A-13B show an example of a modulated pattern sent to aphotocathode (FIG. 13A) and imaged at the patient targeted tissue (FIG.13B). In addition to being steered to the targeted tissue, the patternmay be magnified as needed to cover the appropriate treatment area. Inthis example, the pattern has a magnification of 100 in both planes. Ascan be seen, there may be variations in the pattern that occur duringsteering and/or magnification due to non-linearities in the electronbeam optical components. These variations may be compensated for in theelectron optics, during steering, with the focusing optics, or byadjusting the original pattern on the photocathode. The original patternmay be selected and/or altered so that the resulting electron intensityprofile at the targeted tissue after steering and/or magnification isthe desired electron intensity profile.

C. Photon Collimation System

In accordance with some embodiments of the invention, the system mayutilize a collimation system configured for rapid generation anddelivery of transversely modulated photon beams to targeted tissue forradiation therapy. In certain embodiments, the device includes ascanning pencil-array collimated high-speed intensity-modulated X-raysource (SPHINX). Such a device may include a high efficiency collimationsystem for high energy X-rays produced in medical accelerator. Incontrast to conventional collimators, this collimation system allowsreduced x-ray irradiation times when combined with an electron beam ofsufficient current because the irradiation speed is not limited by slowmechanical motion, thus allowing the ultrafast delivery of radiation tothe targeted tissue desired in a rapid radiation treatment system,including those described therein.

In certain aspects, the precision transverse distribution of the photonsis generated by rastering electron beam pulses from a linear acceleratoronto an array of photon production targets or sheet-shaped target. Bycontrolling the intensity or number of individual electron bunches ateach position, the transverse intensity distribution can be shaped asneeded for treatment.

In accordance with some embodiments, the system operates as follows:electron bunches from a high repetition rate electron accelerator arerastered onto an array of photon production targets. By controlling theindividual electron bunch intensity the produced transverse photondistribution from the array target can be modulated as required by thetreatment. Because the electron bunch intensities and positions can bevaried electromagnetically, the transverse photon beam shaping does notrequire movement of mechanically moving parts, thereby enabling farfaster treatment times than in conventional radiation treatments.

In order to achieve the sub-second treatment times the array photonsource can be combined with an improved high repetition rate linearaccelerator and an improved target/collimator design for photonproduction. In addition, the forward photon intensity characteristics ofthe photon beam potentially can be improved over the traditional high-Ztarget photon production schemes by several strategies including usingdifferent target materials (e.g. lower-Z materials to favor forwardproduction) and different target/collimation geometries (e.g. pintargets, hole targets, cone targets, glancing beam photon production, aswould be known to one of skill in the art, or other strategies asdescribed below in the example variations.

According to certain aspects, it is desirable that the 2-D dosedistribution produced by the array of pencil beams from any givendirection not have dosimetric gaps owing to the space between beams. Asshown for example in FIG. 18A, the space between collimator channels 31should be sufficiently small that the penumbra of individual beams issufficient to fill the gaps between beams at the targeted tissuelocation 20. Conversely, the penumbra should be sufficiently sharp thatthere is good contrast and resolution in the intensity-modulationpattern. Both goals can be fulfilled simultaneously by optimizing thespacing and wall thickness between channels and the length of thechannels. The spacing between channels may allow for overlap between theresultant beamlets, perfect matching, or gaps between beamlets. Anadvantage of overlapping beamlets is more choices for placement ofindividual beamlets for finer targeting, and better field uniformitywhen using multiple adjacent beamlets. The desired intensity patternwill range from uniform coverage of a large area to very highlymodulated patterns with high spatial resolution and contrast.

FIGS. 14A-14B illustrate schematics of such a source in an exampleembodiment, showing the collimator block 30 with its upstream surfaceabutting a bremsstrahlung target 34 and an array of channels 31extending through the block to collimate the photon beams produced bythe bremsstrahlung target 34. For easier visualization in the figure,the spacing between holes is not to scale and represented larger thanwould be used for this purpose. In one aspect, such as shown in FIG. 14Athe channels 31 are circular in cross-section and extend from an inlethole 32 at the upstream surface of the collimator block 30 and extend toan outlet hole 33 at the downstream surface of the collimator block 35,the outlet hole 33 being larger than the inlet hole 32. In anotheraspect, the channels 31 are square or rectangular in cross-section, suchas shown in FIG. 14B and described further in later examples, thechannels extending from an inlet hole 32 at the upstream surface of thecollimator block 30 and extend to an outlet hole 33 at the downstreamsurface of the collimator block 32, the outlet hole 33 being larger thanthe inlet hole 32. In some embodiments, each channel 21 tapers from thesmaller inlet hole 32 to the larger outlet hole 33. While in thisembodiment, the channels generally taper at a fixed rate, in otherembodiments, the taper may of a variable rate or the channels may have afixed diameter along certain portions and taper in certain otherportions of the block.

In one aspect, this source design is useful for a photon source in aradiation treatment system in accordance with the present invention. Theenhanced ability is particularly advantageous for sub-second treatmenttimes for radiation therapy for cancer and other diseases. It isappreciated that the underlying technology for enhanced forward photonproduction can also be used for various radiation treatment, includingthose with longer treatment times, as well as various non-medicalapplications (e.g. truck/cargo scanning for security applications).

Among the advantages of the design described above is rapid andprecisely shaped dose delivery as compared to existing technologies. Inaddition, this design allows for dose modulating and shaping withoutrequiring movement of mechanically moving parts. The source/targetcharacteristics are further improved, as compared to conventionaltreatments, so as to allow further reductions in dosage delivery times,including sub-second dose delivery for treatment of cancer and otherdiseases. This design allows for improved collimation characteristicsand does not require use with multi-leaf collimators that requiremechanical motion to intensity modulate the radiation.

Another advantage of SPHINX compared to MLCs is improved compatibilitywith external magnetic fields such as those used in MRI-guided radiationtherapy systems, because of lack of electric motors and reduced magneticmaterials while still maintaining the ability to intensity modulate thephoton beam and perform dynamic target tracking with the photon beam(e.g., for targets in the heart or other rapidly moving targets).

It is appreciated that the collimator configuration described above maybe varied in a number of ways in keeping with the principles of theinvention, including the following examples:

a. The collimator array consists of channels whose central axes divergefrom the source to the targeted tissue, in order to cover large-volumetargets in the patient (see for example FIGS. 18A, 18C and 18D);

b. The collimator array consists of channels 31 whose central axes 31′are parallel or converge from the source to the targeted tissue 20 (seefor example FIG. 18B), which is more suitable for coveringsmaller-volume targets in the patient;

c. The collimator array has channels 31 with a square profile (see forexample FIG. 14B);

d. The collimator array has channels with a non-square profile that mayhave more optimal area packing (eg, hexagonal, circular, etc.) or edgeenhancement characteristics (see for example FIG. 14A);

e. Rather than an array of channels, the collimator array can be formedby an array of slits (diverging, parallel, or converging) in onedirection in tandem with an array of slits in the perpendiculardirection. A potential advantage of this design is greater ease ofmanufacture than an array of pencil beam channels, at the cost ofgreater space requirement;

f. Different channel sizes may be incorporated into the same array inorder to improve photon throughput, with more photons being deliveredusing larger channels and finer shaping using smaller channels, such asshown for example in FIG. 18C. Alternatively, a similar effect can beachieved through a combination of arrays (e.g., one per treatment headin a multi-beamline system) with each array having a different channelsize and configuration. Alternatively, on a given treatment head, achoice of collimator arrays of different channel configurations may beused (e.g., by mounting them on a carousel to allow switching betweenarrays). The choice of array or individual channels used in the plan canbe made through treatment planning optimization;

g. The channel configuration within a given array may be more complex,for example, consisting of clusters of converging channels but with theclusters diverging from each other, such as shown for example in FIG.18B. Such an arrangement would allow large field coverage with improveddepth dose characteristics;

h. Strategies for heat dissipation include active target cooling,multiple rapid rescanning of the electron beam over the array, andincorporation of appropriately configured heat conducting materials. Insome embodiments, such as shown in the example of FIG. 19A, thebremsstrahlung target may be formed by an array of plugs 34′ of targetmaterial, such as tungsten or other suitable material, aligned with thecollimator channels 31 in the collimator block 35, the array of plugsbeing embedded in a layer heat conducting material 36, such as copper orother suitable material. In addition, the collimator array itself canserve as a heat sink for the target if they are abutting.

i. Strategies for increased forward photon production include applyingelectric and/or magnetic fields to refocus the electron beam within thebremsstrahlung target and choice of target and wall materials, anddividing the target material into multiple thin layers 34″ along thebeam direction to allow electron refocusing between layers, such asshown in the example of FIG. 19B. It is understood that a separationbetween target layers may be fixed or variable to optimize therefocusing effect. A subset of these aspects can be further understoodby referring to the article: W. Ulmer, “On the creation of high energybremsstrahlung and intensity by a multitarget and repeated focusing ofthe scattered electrons by small-angle backscatter at the wall of a coneand magnetic fields—A possible way to improve linear accelerators inradiotherapy and to verify Heisenberg-Euler scatter,” Radiation Physicsand Chemistry 81 (20120 387-402), the entire contents of which areincorporated herein for all purposes.

j. The channels may traverse through the block material at obliqueangles so that the collimator array may be mounted at an angle to thebeamline for more compactness in certain embodiments, such as shown inFIG. 18D-18E;

k. The transmission bremsstrahlung target will be interposed between theelectron beam and the collimator array. Typically, it would consist of alayer of high atomic number such as tungsten although a range ofmaterials may be appropriate. The thickness of the target may be uniformor variable according to an optimized design.

In one aspect, a system includes one or more treatment heads, eachhaving a suitable collimation assemblies disposed within for use in aradiation treatment of a targeted tissue in a patient. In someembodiments, the one or more treatment heads are coupleable with any ofa set of collimator assemblies having different shapes and/orgeometries, such as any of those described therein, which are selectedby a user as desired for a given treatment. In some embodiments, thesystem includes a rotating carousel having multiple differingcollimating assemblies such that selection of a particular collimatingassembly can be effected by rotation of the carousel. In another aspect,the collimating assemblies can be removable from the treatment headssuch that the desired collimating assemblies are selected and attachedto the treatment heads in preparation for the procedure.

In another aspect, one or more beamlines with one or more correspondingcollimator arrays may be on a rotating gantry to create a larger numberof beam directions as illustrated in FIG. 18E, which for simplicityshows 2 beamlines with two beam directions or alternately one beamlinerotating between two directions.

In one aspect, the system may include a rotating gantry 50 having one ormore beamlines and collimation assemblies thereon so as to allowcollimated beams from multiple differing directions. In someembodiments, a rotating carousel may be used include one or morecollimation assemblies or multiple collimation assemblies with differinggeometries. FIG. 18F shows an example of coplanar beamlinesperpendicular to the patient, in which there are multiple beamlines, orin the alternative, FIG. 18F may represent a single or a few beamlinesand a rotating gantry supporting multiple beamlines and collimationassemblies affixed thereon. Of note, a collimation assembly, such as theSPHINX collimation structure may, be used together with conventionalenergy electron linacs (up to ˜20 MeV). In some embodiments, thebeamlines may utilize transport/magnifying optics rather than a SPHINXcollimation structure. Rotation of the gantry allows the beamlines andcollimation assemblies to direct beams toward the targeted tissue fromdifferent directions as needed. In some embodiments, rotation of thegantry would be a requirement if the number of beamlines is small (forexample, less than seven beamlines). In addition, these aspects mayapply to a system in which the beamlines are noncoplanar and at anoblique angle to the patient axis.

FIG. 18G illustrates a schematic of an example system where a singleelectron linear accelerator feeds multiple beamlines (four beamlines inthis example) in a multi-beamline structure 60. A beam deflecting devicemay be configured to move the beam between the different beamlines. Eachbeamline may be used with a collimation assembly, such as a SPHINXstructure, (only one being shown in FIG. 18G) if the appropriate energyrange, or the beamlines may be used without a collimation assembly suchas in a very high energy electron beam treatment.

In one aspect, a system includes one or more treatment heads, eachhaving a suitable collimation assemblies disposed within for use in aradiation treatment of a targeted tissue in a patient. In someembodiments, the one or more treatment heads are coupleable with any ofa set of collimator assemblies having different shapes and/orgeometries, such as any of those described therein, which are selectedby a user as desired for a given treatment. In some embodiments, thesystem includes a rotating gantry having multiple differing collimatingassemblies such that selection of a particular collimating assembly canbe effected by rotation of the gantry. In another aspect, thecollimating assemblies can be removable from the treatment heads suchthat the desired collimating assemblies are selected and attached to thetreatment heads in preparation for the procedure.

Another example geometry for a SPHINX collimation assembly is shown inFIG. 15. In this case, the electron beam impinges on a tungsten target34 of 1.5 mm thickness to produce bremsstrahlung x-rays. Immediatelydownstream of the bremsstrahlung target is the collimator array,consisting of a 10 cm thick block of tungsten with diverging channels.The channel size is such that each forms a photon beamlet of 3 mmdiameter at a distance of 65 cm from the downstream face of the SPHINX,and the spacing is such that the beamlets are 1.5 mm apart (center tocenter) at that distance, and thus fully overlap by a half-width whenadjacent beamlets are formed. This configuration allows finely spacedbeamlet selection as well as no dosimetric gap.

FIG. 15 shows a cross-sectional view along the channels at the upstreamside and at the downstream side of the collimator 30. The projectedwidth of the photon beamlet at the plane of the targeted tissue dependssubstantially on the thickness of the collimator block and the diameterof the outlet hole, which is chosen accordingly. The size of the inlethole should be dimensioned sufficiently large enough to accept the sizeof the electron beam that forms the x-ray source in the bremsstrahlungtarget, and would generally be smaller than the size of the outlet hole.For example, in the embodiment shown, each channel 31 has an inlet hole32 with a half-width of 0.0075 cm and an outlet hole 33 with ahalf-width of 0.02 cm. This results in a beamlet width at the plane ofthe targeted tissue (65 cm downstream of the downstream face of thecollimator) of approximately 3 mm in this example. In one aspect, thecenter to center spacing distance between outlet holes is larger thaneach outlet hole diameter, so that there is sufficient wall thickness toprovide adequate collimation of the x-ray beamlets. The center-to-centerspacing between channels determines the beamlet spacing at the plane ofthe targeted tissue. Diverging, parallel, or converging beamlets areproduced if the spacing at the downstream surface is larger, equal, orsmaller than at the upstream surface. In this embodiment, thecenter-to-center spacing distance between exit holes is 0.06875 cm,while each outlet hole width is 0.04 cm. At the upstream surface of thecollimator 30, along which the bremsstrahlung target 34 is positioned,in this example embodiment, the center-to-center spacing is 0.05625 cmbetween inlet holes, each inlet hole having a width of 0.015 cm. Thisproduces a set of diverging beamlets with 3 mm width and 1.5 mm spacingin the plane of the targeted tissue, and a virtual focal point 120 cmupstream of the plane of the targeted tissue in this example.

FIG. 16 shows the photon fluence when one channel is illuminated by anelectron beam of 10 MeV energy (the bottom channel on the figure) asdetermined by FLUKA Monte Carlo simulation visualized with FLAIR. (SeeFLUKA: a multi-particle transport code” A. Fasso', A. Ferrari, J. Ranft,and P. R. Sala, CERN-2005-10 (2005), INFN/TC_05/11, SLAC-R-773; “FLUKA:a multi-particle transport code,” A. Fasso', A. Ferrari, J. Ranft, andP. R. Sala, CERN-2005-10 (2005), INFN/TC_05/11, SLAC-R-773; and FLAIR,V. Vlachoudis “FLAIR: A Powerful But User Friendly Graphical InterfaceFor FLUKA” Proc. Int. Conf. on Mathematics, Computational Methods &Reactor Physics (M&C 2009), Saratoga Springs, N.Y., 2009; each of whichis incorporated herein by reference in its entirety).

FIG. 17 shows the corresponding dose distribution at 1.5 cm depth in awater phantom. In the example demonstrated, there is a sharp doseprofile of 3 mm width and 2 mm penumbra. A small amount of cross-channelleakage is evident as side wings on the dose distribution with anintensity of approximately 1% of the central peak dose. This isapproximately equal to the level of transmission through a typical MLCleaf but over a much smaller portion of the field. This demonstratesthat the SPHINX design is able to produce spatially modulated beams witha higher degree of modulation and less leakage than MLCs, yet with nomechanical moving parts.

D. General

Numerous specific details are set forth herein to provide a thoroughunderstanding of the claimed subject matter. However, those skilled inthe art will understand that the claimed subject matter may be practicedwithout these specific details. In other instances, methods, apparatusesor systems that would be known by one of ordinary skill have not beendescribed in detail so as not to obscure subject matter that may beclaimed. Items in figures are not drawn to scale, unless otherwiseindicated. When scale is indicated in drawings, the scale may illustratethe advantages of the invention in allowing a more compact system. It isunderstood, however, that the embodiments are not confined to thedimensions shown unless otherwise indicated.

While the present subject matter has been described in detail withrespect to specific embodiments thereof, it will be appreciated thatthose skilled in the art, upon attaining an understanding of theforegoing may readily produce alterations to, variations of, andequivalents to such embodiments. Accordingly, it should be understoodthat the present disclosure has been presented for purposes of examplerather than limitation, and does not preclude inclusion of suchmodifications, variations and/or additions to the present subject matteras would be readily apparent to one of ordinary skill in the art

What is claimed is:
 1. A method for treating a patient, comprising:generating one or more patterned particle beams, wherein each of the oneor more patterned particle beams covers an area of the targeted tissuewith spatially varying beam intensity according to a treatment patternof desired radiation dose distribution; accelerating the one or morepatterned particle beams with one or more accelerators; and transportingand/or magnifying the patterned beam to a desired location, direction,and size suitable for coverage of the targeted tissue, wherein a shape,resolution and contrast of the pattern is suitably maintained duringtransport and/or magnification so as to deliver the desired radiationdose distribution to the targeted tissue according to the treatmentpattern.
 2. The method of claim 1, wherein magnifying the patterned beamcomprises magnifying the beam through one or more focusing elementsdisposed within the beamline of the one or more beams.
 3. The method ofclaim 1, further comprising: steering the one or more patterned particlebeams to the targeted tissue with one or more beam steering devices fromone or more directions.
 4. The method of claim 1, wherein each of theone or more patterned particle beams comprise an array of smallerpatterned beams.
 5. The method of claim 1, wherein the patterned beam ismagnified by up to 100 to 200 times an original size of the pattern ofthe one or more particle beams generated.
 6. The method of claim 5,wherein a resolution of the treatment pattern at the original size iswithin 1/10 of a width of the overall pattern or smaller.
 7. The methodof claim 4, wherein the array of smaller patterned beams is produced byraster scanning individual smaller patterned beams from each beamdirection of the one or more directions.
 8. The method of claim 7,further comprising: forming the two-dimensional intensity-modulatedelectron pattern on a photocathode by projecting or scanning a lightsource onto the photocathode.
 9. A system for treating a patient,comprising: one or more beam generation devices configured to generateone or more patterned particle beams, wherein each of the one or moreparticle beams covers an area of the targeted tissue with spatiallyvarying beam intensity according to a treatment pattern of desiredradiation dose distribution; one or more accelerators configured foraccelerating the one or more patterned particle beams; and one or moremagnification lenses along a beam line of the one or more particle beamsbetween the one or more accelerators and the targeted tissue formagnification of the patterned particle beam to a desired size suitablefor coverage of the targeted tissue according to the treatment pattern.10. The system of claim 9, further comprising: one or more beam steeringdevices configured for steering the one or more patterned particle beamsto the targeted tissue from one or more directions.
 11. The system ofclaim 10, wherein the system is configured such that steering of the oneor more patterned particle beams is performed concurrent with magnifyingthe respective one or more pattern particle beams.
 12. The system ofclaim 9, wherein the one or more magnification lenses comprise aplurality of lenses.
 13. The system of claim 12, wherein a firstmagnetic lenses of the plurality along the beamline has a substantiallyhigher strength than subsequent magnetic lenses of the plurality. 14.The system of claim 9, further comprising: a beam deflector disposedbetween beamlines of the one or more particle beams between the one ormore accelerators and the targeted tissue such that the one or more beamlines can be directed to the targeted tissue from multiple differingangles using a single common accelerator.
 15. The system of claim 9,wherein the system is configured such that the one or more patternedparticle beams comprises an array of smaller patterned beams.
 16. Thesystem of claim 15, wherein the system is configured such that the arrayof smaller patterned beams is produced by raster scanning individualsmaller patterned beams from each beam direction.
 17. The system ofclaim 9, further comprising: an RF powered or DC particle gun and aphoto-cathode configured to produce the two-dimensionalintensity-modulated electron pattern.
 18. The system of claim 9, whereinthe system is configured such that the one or more patterned beamscomprise multiple patterned beams so that, when delivered from multipledirections, the multiple beams produce a desired three-dimensional dosedistribution when summed across the multiple beam directions.
 19. Thesystem of claim 9, further comprising: a controller configured tocontrol delivery of the one or more particle beams to the targetedtissue from the one or more directions thereby irradiating the targetedtissue to deliver an entire treatment dose in less than 10 seconds. 20.The system of claim 19, wherein the controller is configured to rapidlyswitch a modulation pattern sent to the photocathode within a rate ofone pattern every 2 seconds or faster so as to provide delivery ofdiffering treatment patterns to the targeted tissue from multipledirections within less than 10 seconds.
 21. The system of claim 9,wherein the treatment pattern is adapted so as to be suitable for use ina non-medical application.